Embodiments of the invention relate generally to diagnostic imaging and, more particularly, to a method and apparatus for adaptive scatter correction in an imaging system.
Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” shall include anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the subject. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. In typical single energy applications, X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts x-rays to light energy. Each scintillator discharges light energy to a photodiode adjacent thereto, and each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
A CT imaging system may also include an energy sensitive (ES), multi-energy (ME), and/or dual-energy (DE) CT imaging system that may be referred to as an ESCT, MECT, and/or DECT imaging system, in order to acquire data for material decomposition or effective Z or monochromatic image estimation using multiple energy spectra. ESCT/MECT/DECT provides energy discrimination. For example, in the absence of object scatter, the system derives the material attenuation at a different energy based on the signal from two relative regions of photon energy from the spectrum: the low-energy and the high-energy portions of the incident x-ray spectrum.
In a given energy region relevant to medical CT, two physical processes dominate the x-ray attenuation: (1) Compton scatter and the (2) photoelectric effect. These two processes are sensitive to the photon energy and hence each of the atomic elements has a unique energy sensitive attenuation signature. Therefore, the detected signals from two energy regions provide sufficient information to resolve the energy dependence of the material being imaged in order to enable material decomposition. Such systems may use a direct conversion detector material in lieu of a scintillator. Or, in an alternative, a conventional scintillator-based third-generation CT system may be used to provide energy separation measurements by acquiring projections sequentially at different peak kilovoltage (kVp) operating levels of the x-ray tube, which changes the peak and spectrum of energy of the incident photons comprising the emitted x-ray beams. A principle objective of scanning with two distinctive energy spectra (i.e., dual energy) is to obtain diagnostic CT images that enhance information (contrast separation, material specificity, etc.) within the image by utilizing two scans at different polychromatic energy states.
CT systems having an amount of z-coverage that is equal or less than 10 mm (at isocenter), for instance, typically do not use a scatter correction algorithm. However, in recent years CT systems have increasing z-coverage in order to shorten scan times and reduce overall dose. The goal has been to obtain an image of an object, such as a cardiac region, in a single rotation. As CT systems have grown in z-coverage (i.e., increased numbers of slices), however, scatter has become an increasingly significant factor. For example, for a 16-slice scanner with 10 mm z-coverage, the scatter-to-primary ratio (SPR) is less than 10% for a 35 cm poly phantom. When the z-coverage increases to 40 mm (or 64 slices), the SPR increases to 20% for the same size phantom. And, in a 160 mm wide-cone system, SPR can reach 28% for large objects and a typical 1D anti-scatter grid, or 8% with a 2D anti-scatter grid. It is well-known that an increased SPR degrades image quality due to image artifact and contrast loss.
The amount of scatter in a CT system also depends in part on an amount of energy in the projection beam. Thus, for lower energy applications, below 80 kVp for instance, SPR is greater than for higher energy applications, further exacerbating the issue of scatter and the ability to correct for it. Thus, inherent in a dual energy application, scatter (particularly at the low kVp operation of a dual energy procedure) correction may be necessary, moreso for the low kVp data of such an operation.
Many attempts have been made in the past to improve the scatter performance of CT systems. For example, hardware improvements may be implemented by increasing the aspect ratio of post-patient collimation plates, significantly improving the amount of scatter rejection. The aspect ratio for a collimator is typically defined as the collimator plate height (H) divided by the aperture width (W). In general, the higher the aspect ratio, the better is the scatter rejection capability. However, such solutions tend to be expensive, may limit performance, and may increase an amount of dose required to obtain adequate image data. Thus, in addition to hardware solutions to limit scatter, scatter correction methods have been developed as wider coverage CT systems have been developed.
Generally, there are two types of scatter correction for cone-beam CT: 1) direct correction in x-ray projection space, or 2) using a second pass algorithm using reconstructed images.
When correcting in projection space, known solutions include estimating scatter in projection data using an empirical function (such as a square root function). Such solutions may be computationally attractive, but may have limitations that are exacerbated in wider z-coverage and low energy applications. Typically, such solutions are not accurate for non-uniform objects. Also, assuming a constant empirical function correction may not accurately predict scatter intensity across various scan conditions, and in some applications the amplitude of an estimated scatter profile has to be reduced in order to meet image quality requirements. Further, estimations assuming a constant correction value do not take into account scatter related to a bowtie filter.
When correcting in image space, some known solutions for scatter correction include estimating an amount of scatter based on images and using, for instance, a Monte Carlo application. In this approach, scatter profiles are computed by tracking rays through a reconstructed volume. One known solution includes estimating a size of an image and then performing a scatter correction based on the estimation. However, correcting in image space takes much longer computational time when compared to the first approach, above, which is compounded as wider coverage systems are developed.
Therefore, it would be desirable to design an apparatus and method for improving scatter correction for CT imaging.